Detectors comprising a one-dimensional or two-dimensional array of detector pixels are commonly used for radiation imaging, for example, in medical applications such as nuclear medicine, radiology, radiotherapy or proton-therapy. In a radiation detector array, each pixel may be adapted for generating a signal, such as an electric current, an electric charge or an electric potential, that is indicative for a dosimetric or radiative quantity related to radiation incident on that pixel. For example, silicon diodes and ionization chambers may be commonly used in pixel elements for radiation detection applications.
Pixel signals may be read out in parallel using a front-end electronic subsystem comprising a plurality of multi-channel integrated circuits. For instance, in an imaging array known in the field, the IBA MatriXX detector array, the current from 1024 ionization chambers may be integrated in parallel by 16 dedicated chips, each having 64 input channels. Thus, such imaging system may also require an efficient means for routing and switching of the signals from the pixels to the readout channels of the front-end electronic subsystem. This signal routing may affect the dynamic range and the image quality of the images produced by the device, and may furthermore influence the achievable detector area and pixel pitch due to limitations in available area for routing.
At least one interposer printed circuit board (PCB) may connect the sensor pixels to the readout chips. For example, all signal lines may be provided on a PCB which may also form an integral part of the sensor array, while the readout chips may be connected to these signal lines, e.g. using a connector mechanism to allow easy replacement and/or maintenance. For example, each readout chip may be soldered onto a small carrier PCB that can be plugged into the interposer PCB using connectors.
United States patent application US 2005/286682 discloses a detector for use in a radiation imaging system. This detector module comprises a sensor array for converting X-ray signals to electrical signals, at least one electronic device for converting the electrical signals to a corresponding digital signal, and a switching circuit for routing the electrical signals from the sensor array to the electronic device.
X-ray detectors for imaging applications are known, e.g., from US 2004/0136493 and US 2005/0173642. These detectors comprise a 2D-array of pixels. These pixels are organised as n lines by m columns and require a switched element, such as an FET, at each pixel. The individual pixels are accessed by sequentially addressing the n scan lines and acquiring the pixels values from the m pixels of each line through the m data lines. Such designs have the advantage that only m channels of readout circuits are necessary. The switching elements may introduce measurement errors, which may be not acceptable, especially when weak signal values are expected. In addition, when the detector is to be used in high dose applications, such as radiotherapy machine QA, or on-line or off-line treatment plan verification, the switching element may not be resistant to the high radiation dose.
A pixelated photon counting mode detector is known from US 2005/139757. In this detector, the readout circuits, implemented in an ASIC chip, are connected to a detector array through a ball grid array package having a plurality of solder balls formed on a first side and a plurality of electrical contacts formed on a second side. The readout circuits are within each pixel's geometric area. The readout circuits are located in immediate vicinity to the corresponding detectors, and submitted to the same radiation field.
The routing of the pixel signals to the front-end electronics can be a complicated task. The routing between the array pixels and the input channels of the readout chips may be optimized to simplify the layout and/or to obtain a good performance. For example, parameters that may be tuned by such optimization of the design are the length of the signal lines, the line impedance, the number of via connections, the cross talk between channels, the sensitivity to electromagnetic disturbances and the radiation of electromagnetic disturbances.
However, the overall cost of readout electronics may represent a large fraction of the overall detector price, particularly for detectors with many pixels. State-of-the-art readout chips may be quite expensive, and may be required to be mounted on advanced PCBs as well. For example, in a Stereotactic Body Radiation Therapy (SBRT) imaging array comprising a large number of silicon diode pixels, e.g. about 2000 pixels or more, the cost of readout electronics could amount to about 40% of the production cost.
It is particularly advantageous to achieve a high spatial resolution and a small pixel pitch in a detector array. For example, two-dimensional detectors having a small pixel pitch can be used for accurately measuring radiation fields characterized by high radiation dose gradients, e.g. for stereotactic radiotherapy applications. However, for two-dimensional detector arrays covering a predetermined area, the number of pixels is proportional to (1/pitch)2. The cost and design complexity may therefore also increase in proportion to (1/pitch)2. This implies that an optimal trade-off may be sought between spatial resolution and device complexity, e.g. as determined by the cost versus the benefit to the user.
As an example, if a 12×12 cm2 active area is covered with 3 mm pixel pitch, 1600 channels may be needed. However, this number rises to 2304 when the pitch is only slightly reduced to 2.5 mm. The cost of a pixel array may depend mainly on the area, e.g. for technologies such as ionization chamber arrays implemented on a PCB or monolithic silicon, while the cost of the corresponding readout electronics rises with the number of pixels.